1. Field of the Invention
The present invention relates to a reinforced responsive polymeric system. More specifically, the present invention relates to a polymeric system comprising at least two phases, an environmentally responsive polymeric component and a solid reinforcing component, which is introducible into the body and which undergoes a change in viscosity at a predetermined body site, said polymeric system being useful in drug delivery, tissue engineering, gene therapy and other biomedical applications.
2. Prior Art
There is a wide variety of materials which are foreign to the human body and which are used in direct contact with its organs, tissues and fluids. These materials are called Biomaterials, and they include, among others, polymers, ceramics, biological materials, metals, composite materials and combinations thereof.
The development of polymers suitable to be implanted without requiring a surgical procedure, usually named injectable polymers, has triggered much attention in recent years. These materials combine low viscosity at the injection stage, with a gel or solid consistency developed in situ, later on. The systems of the present invention are preferably used, without limitation, as matrices for the controlled release of biologically active agents, as sealants, as coatings and as barriers in the body. The area of Tissue Engineering represents an additional important field of application of the reinforced responsive systems disclosed hereby, where they can perform as the matrix for cell growth and tissue scaffolding.
The syringeability of injectable biomedical systems is their most essential advantage, since it allows their introduction into the body using minimally invasive techniques. Furthermore, their low viscosity and substantial flowability at the insertion time, enable them to reach and fill spaces, otherwise inaccessible, as well as to achieve enhanced attachment and improved conformability to the tissues at the implantation site. On the other hand, the sharp increase in viscosity is a fundamental requirement for these materials to be able to fulfill any physical or mechanical function, such as sealing or performing as a barrier between tissue planes. The high viscosities attained play also a critical role in generating syringable materials that, once at the implantation site, are also able to control the rate of release of drugs or can function as the matrix for cell growth and tissue scaffolding. Clearly, biodegradability is yet another important requirement for some of these materials.
A polymer network is characterized by the positive molecular interactions existing between the different components of the system. These inter-reactions may be physical in nature, such as chain entanglements, or chemical such as ionic interactions, hydrogen bonding, Van der Waals attractions and covalent bonding. Bromberg et al. (U.S. Pat. No. 5,939,485) developed responsive polymer networks exhibiting the property of reversible gelation triggered by a change in diverse environmental stimuli, such as temperature, pH and ionic strength. Pathak et al. (U.S. Pat. No. 6,201,065) disclosed thermo-responsive macromers based on cross-linkable polyols, such as PEO-PPO-PEO triblocks, capable of gelling in an aqueous solution. The macromers can be covalently crosslinked to form a gel on a tissue surface in vivo. The gels are useful in a variety of medical applications including drug delivery.
Park et al. (U.S. Pat. No. 6,271,278) developed superporous hydrogel composites formed by polymerizing one or more ethylenically-unsaturated monomers, and a multiolefinic cross-linking agent, in the presence of particles of a disintegrant and a blowing agent.
The term “thermosensitive” refers to the capability of a polymeric system to achieve significant chemical, mechanical or physical changes due to small temperature differentials. The resulting change is based on different mechanisms such as ionization and entropy gain due to water molecules release, among others (Alexandridis and Hatton, Colloids and Surfaces A, 96, 1 (1995)). Since one of their fundamental advantages is to avoid the need for an open surgical procedure, thermo-responsive materials are required to be easily syringable, combining low viscosity at the injection stage, with a gel or solid consistency being developed later on, in situ.
Thermosensitive gels can be classified into two categories: (a) if they have an upper critical solution temperature (UCST), they are named positive-sensitive hydrogels and they contract upon cooling below the UCST, or (b) if they have a lower critical solution temperature (LCST), the are called negative-sensitive hydrogels and they contract upon heating above this temperature.
The reverse thermo-responsive phenomenon is usually known as Reversed Thermal Gelation (RTG) and it constitutes one of the most promising strategies for the development of injectable systems. The water solutions of these materials display low viscosity at ambient temperature, and exhibit a sharp viscosity increase as temperature rises within a very narrow temperature interval, producing a semi-solid gel once they reach body temperature. There are several RTG displaying polymers. Among them, poly(N-isopropyl acrylamide) (PNIPAAm) (Tanaka and co-workers in U.S. Pat. No. 5,403,893 and Hoffman A. S. et al., J. Biomed. Mater. Res., 24, 21 (1990)), PEG-PLGA-PEG triblock polymers (Jeong et al., Nature, 388, 860 (1997)), etc. Unfortunately, poly(N-isopropyl acrylamide) is non-degradable and, in consequence, is not suitable for a diversity of applications where biodegradability is required.
Definitely one of the most important RTG-displaying materials is the family of poly(ethylene oxide)/poly(propylene oxide)/poly(ethylene oxide) (PEO-PPO-PEO) triblocks, available commercially as PluronicRTM (Krezanoski in U.S. Pat. No. 4,188,373). Adjusting the concentration of the polymer, renders the solution with the desired liquid-gel transition. However, relatively high concentrations of the triblock are required (typically above 15-20%) are required to produce compositions that exhibit such a transition, even minor, at commercially or physiologically useful temperatures. Another known system which is liquid at room temperature, and becomes a semi-solid when warmed to about body temperature, is disclosed in U.S. Pat. No. 5,252,318, and consists of tetrafunctional block polymers of polyoxyethylene and polyoxypropylene condensed with ethylenediamine (commercially available as Tetronic.RTM).
The endothermic phase transition taking place, is driven by the entropy gain caused by the release of water molecules bound to the hydrophobic groups in the polymer backbone. Unfortunately, despite of their potential, some fundamental aspects of their performance severely restrict their clinical use. Even though these materials exhibit a significant increase in viscosity when heated up to 37° C., the levels of viscosity attained are not high enough for most clinical applications. Derived from this fundamental limitation, these systems display unsatisfactory mechanical properties and unacceptably short residence times at the implantation site. Furthermore, due to these characteristics, these gels have high permeabilities, a property which renders them unsuitable for drug delivery applications because of the fast drug release kinetics of these gels. Despite of their clinical potential, these materials have failed to be used successfully in the clinic, because of serious performance limitations (Esposito et al., Int. J. Pharm. 142, 9 (1996)).
Composite Materials comprising a matrix and a reinforcing component have recently attracted much attention as new materials for improved biomedical devices. Usually used reinforcing components are polymeric or ceramic materials, as well as metals, carbons (mainly fibers), and biological materials. The reinforcement may be in a fibrous or particulate form, or creating specific three dimensional constructs. Examples of the latter are amorphous lattice structures, meshes and fab, non-woven structures, a filament wound or braided structures, or honeycomb structures. Also, the reinforcing constituent may be a macro, micro or nano-sized material and it may be hollow, porous or solid.
Whereas the vast majority of the biomedical composites are stiff structures, engineered to perform in conjunction with hard tissues, work has also been conducted aiming at developing composite materials for soft tissue implants, such as arterial prostheses, pericardial and hernial patches and tracheal conduits. So far, the major uses of biomedical composites have been in reconstructive surgery for joint replacement, as ligaments and tendons and as a variety of fixation devices. While the matrix is generally a polymer, typically, devices comprised carbon and glass fibers. Further examples of the expanding use of composite materials in diverse orthopaedic and reconstructive applications are glass fibre polyurethane splints and porous PTFE-carbon fiber patches and composites in maxillofacial and skull defect reconstruction. Additional novedous uses, such as intramedular nails and ureter prosthesis are described by S. Ramakrishna et al. (Composites Science and Technology, 61, 1189 (2001)).
The biocompatibility of the prosthesis is a necessary but by no means sufficient condition for a successful implant action. The healing and remodeling of the natural tissue and the successful incorporation of the implant are greatly affected by the stress field induced on the tissue. Because the mechanical properties of the implant have a determining effect on these phenomena, the importance of implants that mimic the mechanical response of the replaced tissue becomes apparent. The main characteristics of this response will vary significantly with the type of tissue, e.g. osseous calcified versus soft connective.
Anisotropy, an inherent merit of fibrous composite materials, is an important characteristic common to most biological tissues. For example, blood vessels are complex, multilayered structures, comprising collagen and elastin fibers, smooth muscle cells, ground substance and endothelium. The anisotropy of blood vessels is because of the orientation of their fibrous components. An additional important characteristic shared by most natural fibrous soft tissues pertains to their non-Hookean dimensional response under physiological loading modes, where J-shaped stress-strain curves are exhibited.
Biodegradability plays a unique role in a diversity of devices, implants and prostheses, being this property an additional important requirement for some of these materials. Their most obvious advantage pertains to the fact that there is no need to remove the system, once it has accomplished its objectives. In addition, they can perform as matrices for the release of bioactive molecules and result in improved healing and tissue regeneration processes. Biodegradable polymers such as polyesters of α-hydroxy acids, like lactic acid or glycolic acid, are used in diverse applications such as bioabsorbable surgical sutures and staples, some orthopedic and dental devices, drug delivery systems and more advanced applications such as the absorbable component of selectively biodegradable vascular grafts, or as the temporary scaffold for tissue engineering. The synthesis and biodegradability of poly(lactic acid) was reported by several groups (Tormala P. and group, Biomaterials, 16, 1353 (1995) and Gopferich A., Biomaterials, 17, 103 (1996)). Biodegradable polyanhydrides (Langer, R., J. Biomed. Mat. Res., 28, 1465 (1994)) and polyorthoesters (Gurny R., ‘Polymer Biomaterials in Solution, as Interfaces and as Solids’, Page 683, S. L. Cooper, C. H. Bamford and T. Tsuruta (Editors), VSP-Utrecht, The Netherlands (1995)) having labile backbone linkages, have been developed, the disclosures of which are incorporated herein. Polymers which degrade into naturally occurring materials, such as polyaminoacids, also have been synthesized. Degradable polymers formed by copolymerization of lactide, glycolide, and ε-caprolactone have been disclosed (Kissel T. and collaborators, J. Biomed. Mater. Res., 30, 31-40 (1996). Polyether-polyester combinations especially of polyethylene glycol (PEG) and aliphatic polyesters like poly(lactic acid), poly(glycolic acid) and poly(caprolactone), either as a blend or as a copolymer, in order to increase the hydrophilicity and degradation rate, have been reported. Most of the work was focused on poly(ethylene glycol)/poly(glycolic) (PEG-PGA) or poly(lactic) (PEG-PLA) acid materials (Cohn et al., Polymer, 28, 2018-2022 (1987) and J. Biomed. Mater. Res., 21, 1301-1316 (1987) and Sawhney S. A. and Hubbell J. A., J. Biomed. Mater. Res. 24, 1397 (1990)). Furthermore, these polymers present relatively fast degradation rates, from a few days to a few months. This drawback constitutes one of the relevant application limitations. Another group of poly(ether-ester)s is the poly(ethylene glycol)-poly(caprolactone) (PEG-PCL)-based polymers. Thus, a broad work was done on high MW PEG-PCL block copolymers. Vert and co-workers (Polym. Int., 45, 419 (1998) synthesized and characterized PEG-PCL copolymers of intermediate molar masses with both PEG and PCL crystallizable blocks, using dicyclohexylcarbodiimide as coupling agent. Findings of cytotoxicity and hemocompatibility tests showed biocompatibility. Lee and partners (J. Control. Release, 73, 315 (2001) reported amphiphilic block copolymeric micellar systems composed of methoxy poly(ethylene glycol)/epsilon-caprolactone for DDS. Cohn et al (J. Biomed. Mater. Res. 59, 273 (2002) produced series of PEG-PCL-containing biodegradable poly(ether-ester-urethane)s, covering a wide range of compositions. Finally, reduction of adhesions associated with post-operative surgery based on the administration of polymeric composition comprising chain-extended poly(hydroxy-carboxylic acid)/poly(oxyalkylene) ABA triblocks to a site in the body which has been subjected to trauma, e.g. by surgery, excision or inflammatory disease was described (Cohn et al. in U.S. Pat. Nos. 5,711,958 and 6,136,333).
Unfortunately, the few absorbable polymers clinically available today are hydrophobic solids which are, therefore, clearly unsuitable for non-invasive surgical procedures, where injectability is a fundamental requirement. The only way to avoid the surgical procedure with these polymers, is to inject them as micro or nanoparticles or capsules, typically containing a drug to be released. As an example, injectable implants comprising calcium phosphate particles in aqueous viscous polymeric gels, were first proposed by Wallace et al. in U.S. Pat. No. 5,204,382. Even though these the ceramic component is generally considered to be nontoxic, the use of nonabsorbable particulate material seems to trigger a foreign body response both at the site of implantation as well as at remote sites, due to the migration of the particles, over time.
Among the approaches developed, the in situ precipitation technique developed by R. Dunn, as disclosed in U.S. Pat. No. 4,938,763, is one strategy worth mentioning. These systems comprise a water soluble organic solvent, in which the polymer is soluble. Once the system is injected, the organic solvent gradually dissolves in the aqueous biological medium, leaving behind an increasingly concentrated polymer solution, until the polymer precipitates, generating the solid implant in situ. A similar approach has been reported by Kost et al (J. Biomed. Mater. Res., 50, 388 (2000)).
In situ polymerization and/or cross-linking is another important technique used to generate injectable polymeric systems. Hubbell et al described in U.S. Pat. No. 5,410,016, water soluble low molecular precursors having at least two polymerizable groups, that are syringed into the site and then polymerized and/or crosslinked in situ chemically or preferably by exposing the system to UV or visible radiation. Mikos et al (Biomaterials, 21, 2405 (2000)) described similar systems, whereas Langer et al (Biomaterials, 21, 259 (2000)) developed injectable polymeric systems based on the percutaneous polymerization of precursors, using UV radiation. An additional approach was disclosed by Scopelianos and co-workers in U.S. Pat. No. 5,824,333 based on the injection of hydrophobic bioabsorbable liquid copolymers, suitable for use in soft tissue repair.
Unfortunately, all these techniques have serious drawbacks and limitations, which significantly restrict their applicability. The paradox in this area has to do, therefore, with the large gap existing between the steadily increasing clinical demand for Injectables, on one hand, and the paucity of materials suitable to address that need, on the other hand.
The emerging field of Tissue Engineering and its nascent clinical application, represents a major breakthrough both conceptually as well as technologically. The objective of Tissue Engineering is to induce regeneration of functional tissue, by providing the appropriate three-dimensional scaffolding construct on which cells will be able to grow, differentiate and generate new tissue. Tissue Engineering systems comprise a matrix containing the cells and a scaffold which functions as a substrate for cells attachment. Customarily, the matrix consists of natural or synthetic hydrogels such as alginates, hyaluronic acid, collagen gels and additional materials such as fibroin. On the other hand, the scaffolds typically comprise biodegradable aliphatic polyesters, such as polylactic acid, polyglycolic acid and polycaprolactone and copolymers (Hutmacher, Biomaterials, 21, 2529-2543 (2000)). Clearly, the composition and mechanical properties of the materials, strongly affect the ability of the system to actively promote the regeneration of autologous functional tissue. In addition, the macrostructural characteristics of the scaffold, play also a fundamental role in determining the type of cells and tissue components present in the new tissue. The template's ultimate task is to provide a gradually disappearing, temporary construct for the generation of viable new tissue. Therefore, if autologous tissue is to regenerate and replace the construct, biodegradability is one of its indispensable attributes. It is also necessary for the template to perform as an adhesive substrate for cells, promoting their growth and differentiation, while retaining cell function. Also, for a scaffold to perform successfully, it is required to be biocompatible, to display the right porosity and to be mechanically suitable.
In general terms, Tissue Engineering can be classified into in vitro and in vivo types. While the former concentrates on the ex vivo generation of tissues from cells removed from a donor site, the latter aims at regenerating functional tissue at the site of implantation, by the combined action of biomolecules and cells, in situ.